Cardiac support systems and methods for chronic use

ABSTRACT

A high efficiency cardiac support system is suitable for chronic use in treating heart failure, wherein the system includes an implantable rotary blood pump, an implantable power module, a wireless power transfer subsystem, a patient monitor, and a programmer. In a cardiac support system, the cumulative efficiencies of the components of the system are capable of providing therapeutically effective blood flow for a typical day of awake hours using the energy from a single wireless recharge of an implanted rechargeable energy source. Moreover, the implantable rechargeable energy source may be recharged during a normal sleep period of 8 hours or less. The system may provide full or partial cardiac support without the need for external wearable batteries, controllers, or cables.

RELATED APPLICATIONS

This application claims the benefit of U.S. patent application Ser. No.12/899,748 to Akkerman et al., filed on Oct. 7, 2010; and U.S.Provisional Patent Application No. 61/421,779 to Armstrong et al., filedon Dec. 10, 2010, which are incorporated herein by reference.

FIELD OF THE INVENTION

This invention relates to implantable cardiac support systems,particularly implantable blood pump systems for the treatment of heartfailure.

BACKGROUND OF INVENTION

In patients with heart failure, there is a need for therapeuticallyenhancing blood flow using an implantable system. Cardiac supportsystems include, but are not limited to, left ventricular assist devices(LVADs), right ventricular assist devices (RVADs), using two devices toassist both ventricles as a bi-ventricular assist device (BiVADs), andtotal artificial hearts (TAHs). Ventricular assist devices (VADs) thatare suitable for adults may call for approximately 5 liters/min (LPM) ofblood flow at 100 mm of Hg differential pressure which equates to about1 watt of hydraulic power. Currently available implantable continuousflow blood pumps consume significantly more electric power to producethe desired amount of flow and pressure.

High pump power consumption of current systems may make it impracticalto implant a power source of sufficient capacity for a full day of awakehours of operation in the body. For example, size restrictions ofimplantable power sources may only allow the implantable power source toprovide up to an hour of operation time. Instead, high power consumptionblood pumps may provide a wire connected to the pump that exits the body(i.e. percutaneous) for connection to a power source that issignificantly larger than an implantable power source. These blood pumpsmay require external power to be provided at all times to operate. Inorder to provide some mobility, external bulky batteries and controllersmay be utilized. However, percutaneous wires and externally worncomponents can still restrict the mobility of a person with such a bloodpump implant. For example, such high power consumption blood pumps haveexternal batteries that frequently require recharging thereby limitingthe amount of time the person can be away from a charger or powersource, external batteries and controllers that can be heavy orburdensome thereby restricting mobility, percutaneous wire skinpenetrations that are not suitable for prolonged exposure to watersubmersion (i.e. swimming, bathing, etc.), and/or other additionaldrawbacks.

For example, negative impacts of these types of systems may includesusceptibility to infection, constraints on sleep position, restrictionson water activities such as swimming and bathing, concern for wireentanglement or severing, necessity to avoid static discharges, and amultitude of others. Furthermore, the external batteries and controlsystems are burdensome. It would be advantageous to eliminate thepercutaneous wire and burdensome external batteries and control system.

While there is limited use of wireless power systems in some neuralstimulators, widespread use of wireless power systems for implantableheart pumps has not been adopted. Currently, few applications ofwireless power transfer have been applied to VADs or TAHs due to thehigher power transfer levels required, relatively high power consumptionof such devices, limited space available for implantable rechargeablebatteries, limited capacity of implantable rechargeable batteries, andthe like.

However, in order to overcome issues associated with percutaneous wires,some wireless power transfer systems have been developed that useinductive coupling between an implanted coil and an external coil totransfer power across the skin, thereby obviating the need for apercutaneous wire. This type of wireless power transfer system simplyuses the inductive effect between two coils similar to a standardtransformer. This approach has been used widely to recharge implantedbatteries in some neural stimulators. Further, these inductive systemsmay require precise alignment between the two coils, and may requireclose spacing between coils on the order of a few inches or less. Theseinductive systems can generate significant amounts of heat near theskin, and require the patient to be immobile during charging if theexternal power source is not easily mobile. Energy lost by such systemsis generally released as heat that is dissipated into the human body,which may produce heat-related health complications or requireadditional components to compensate for the heat generated.

LionHeart LVD-2000 LVAD from Arrow International, Inc. and theHeartSaver™ LVAD from WorldHeart Corporation eliminated the percutaneouswire by powering the implanted portion using inductively-coupledTranscutaneous Energy Transfer (TET). These systems eliminated the wire,but did not eliminate the burdensome external batteries and controlsystem which still had to be worn by the patient. For example, theLionHeart LVD-2000 had a rechargeable implantable battery for briefperiods when the external power was unavailable or needed to be removed.However, due to the energy demands of the implanted system, thatimplantable battery could supply only about 20 minutes of energy. Notethat the size of an acceptable system for implanting into a patientconstrains the capacity of implantable energy storage. Consequently,although the LionHeart LVD-2000 did not require a percutaneous wire, theburden of the external batteries and controller remained similar to thatof systems with a percutaneous wire.

Implantable cardiac support systems have numerous sources of potentialenergy inefficiency. To produce a therapeutically enhanced blood flow,power is needed to produce a particular desired blood flow rate at aparticular desired pressure. Blood flow may be imparted by anelectro-mechanical device, such as by a rotary pump. The design of theelectro-mechanical device is critical to efficiently transferring theelectrical energy powering the device into the desired blood flow.Further, a cardiac support system may also include an energy storagesystem. The design and operation of the energy storage system iscritical to efficiently maintain and transfer stored electrical energy.

As a result of the significant drawbacks of existing systems, there isan unmet need for an energy-efficient cardiac support system capable ofeliminating percutaneous wires for power or control, and doing sowithout burdening the patient with external batteries or controllers.There is an unmet need for a system which not only restores cardiacfunction, but restores an unburdened ambulatory lifestyle.

SUMMARY OF THE INVENTION

In an illustrative implementation, a cardiac support system includes arotary blood pump that is implantable into the human body, wherein therotary blood pump generates a desired amount of blood flow; and a powermodule connected to the rotary blood pump, wherein the power modulestores electrical energy utilized to operate the rotary blood pump, andthe power module is implantable into the human body. The system furtherincludes a receiving coil assembly coupled to the power module, whereinthe receiving coil assembly is implantable into the human body, and atransmitting coil assembly magnetic resonance coupled to the receivingcoil assembly, wherein the transmitting coil assembly is utilized toelectromagnetically transfer energy to the receiving coil assembly.

In another illustrative implementation, a method for providing cardiacsupport to a patient includes generating a desired amount of blood flowwith a rotary blood pump implanted in the patient; storing electricalenergy in a power module implanted in the patient, wherein the powermodule stores electrical energy received from a receiving coil assemblyimplanted in the patient; and coupling a transmitting coil assembly tothe receiving coil assembly using magnetic resonance coupling, whereinthe transmitting coil assembly electromagnetically transfers energy tothe receiving coil assembly.

In yet another illustrative implementation, a cardiac support systemincludes a rotary blood pump that is implantable into the human body,wherein said rotary blood pump generates a desired amount of blood flow,and a power module connected to said rotary blood pump, wherein saidpower module is implantable into said human body. The system alsoincludes a receiving coil assembly receiving energy wirelessly, whereinsaid receiving coil assembly is implantable into said human body, saidreceiving coil assembly transfers said energy into said power module,and said desired amount of blood flow is generated with an EnergyConversion Ratio (ECR) of 1.0 or greater using energy stored by saidpower module. The ECR is defined as the sustained flow rate (in LPM) acardiac support system can provide against 100 mm-Hg differentialpressure for 24 hours from a 40 Watt-hour rechargeable energy source.

In yet another illustrative implementation, a method for treating heartfailure is disclosed. The method includes receiving energy wirelesslyvia a receiving coil assembly, wherein said receiving coil assembly isimplantable into the human body, and storing said energy in a powermodule, wherein said power module is implantable into said human body.The method also includes producing a desired amount of blood flow from arotary blood pump utilizing energy provided by said power module,wherein said rotary blood pump is implantable into said human body, andsaid desired amount of blood flow is generated with an Energy ConversionRatio (ECR) of 1.0 or greater using energy stored by said power module.

The foregoing has outlined rather broadly various features of thepresent disclosure in order that the detailed description that followsmay be better understood. Additional features and advantages of thedisclosure will be described hereinafter.

BRIEF DESCRIPTION OF THE DRAWINGS

For a more complete understanding of the present disclosure, and theadvantages thereof, reference is now made to the following descriptionsto be taken in conjunction with the accompanying drawings describingspecific implementations of the disclosure, wherein:

FIG. 1 is a top view of an illustrative implementation of a rotary bloodpump;

FIG. 2 is a cross-sectional side view of an illustrative implementationof a rotary blood pump;

FIG. 3 is a cross-sectional top view of an illustrative implementationof a rotary blood pump;

FIG. 4 is a close up cross-sectional view of an area of an illustrativeimplementation of a rotary blood pump;

FIG. 5 is a cross-sectional view of an illustrative implementation of animpeller for use in a rotary blood pump;

FIG. 6 is a cross-sectional view of an illustrative implementation of arotary blood pump housing;

FIG. 7 is a cross-sectional view of an illustrative implementation of amotor housing for use in a rotary blood pump;

FIG. 8 is an isometric view of an illustrative implementation of animpeller for use in a rotary blood pump;

FIG. 9A-9K are illustrative implementations of various types of patterngrooves;

FIG. 10A-10D are cross-sectional views of various shapes of patterngrooves;

FIG. 11 is a cross-sectional side view of an illustrative implementationof a rotary blood pump with an axial hydrodynamic bearing;

FIGS. 12A and 12B are top views of illustrative implementations ofimpellers with spiral herringbone grooves and spiral grooves;

FIG. 13 is a close up cross-sectional view of an area of an illustrativeimplementation of a rotary blood pump with an axial hydrodynamicbearing;

FIGS. 14A and 14B are isometric views of illustrative implementation ofimpellers with spiral herringbone grooves and spiral grooves;

FIG. 15 is a cross-sectional side view of an illustrative implementationof a rotary blood pump with a conically shaped impeller;

FIG. 16A-16E are isometric views of illustrative implementations ofconically shaped impellers;

FIG. 17 is a close up cross-sectional view of an area of animplementation of a rotary blood pump with a conically shaped impeller;

FIG. 18 is a cross-sectional side view of an illustrative implementationof a rotary blood pump with passive magnetic axial bearings;

FIG. 19 is a cross-sectional top view of an illustrative implementationof a rotary blood pump with passive magnetic axial bearings;

FIG. 20 is a close up cross-sectional view of an area of an illustrativeimplementation of a rotary blood pump with passive magnetic axialbearings.

FIG. 21 is an illustrative implementation of a wireless power system;

FIG. 22 is an isometric view of an illustrative implementation of atransmitting coil assembly;

FIG. 23 is a front view of an illustrative implementation of atransmitting resonant coil and excitation coil;

FIG. 24 is an isometric view of an illustrative implementation of areceiving coil assembly;

FIG. 25 is a front view of an illustrative implementation of a receivingresonant coil and power pick-up coil;

FIGS. 26 a and 26 b are front and side views of illustrativeimplementations of a resonant coil with single wrap conductive foil;

FIG. 27 is an illustrative implementation of a wireless power systemwith a sympathetic coil;

FIG. 28 is a functional block diagram of a rectifier and DC filtercomponents connected to a power pick-up coil.

FIG. 29 is an illustrative implementation of anatomical positioning ofcomponents of a cardiac support system; and

FIG. 30 is a block diagram of an implantable power module.

DETAILED DESCRIPTION

Refer now to the drawings wherein depicted elements are not necessarilyshown to scale and wherein like or similar elements may be designated bythe same reference numeral through the several views.

Referring to the drawings in general, it will be understood that theillustrations are for the purpose of describing particularimplementations of the disclosure and are not intended to be limitingthereto. While most of the terms used herein will be recognizable tothose of ordinary skill in the art, it should be understood that whennot explicitly defined, terms should be interpreted as adopting ameaning presently accepted by those of ordinary skill in the art.

The following detailed description provides for an implantable, energyefficient, small, and wireless cardiac support system. The cardiacsupport system may include an implantable rotary blood pump, animplantable power module, and a wireless power transfer subsystem. Theimplantable rotary blood pump may be powered by an implantable powermodule which can be recharged using a wireless power transfer subsystem.Those skilled in the art will appreciate that the various featuresdiscussed below can be combined in various manners, in addition to theimplementations discussed below. The scope of the claims is in no waylimited to the specific implementations discussed herein.

In response to unmet needs, the cardiac support system may use apower-efficient rotary blood pump, an efficient power module, and awireless power subsystem, and/or other components, as discussed furtherherein. Currently available cardiac support systems may have severaldrawbacks, such as percutaneous wire(s), bulky external battery packsand controllers, and/or required frequent and close distance recharging.The cardiac support system discussed herein is capable of providing theenhanced blood flow needed by an average patient during a typical day ofawake hours utilizing power stored by an implanted energy storagedevice. The implanted energy storage device is capable of beingwirelessly recharged over a significant charging distance during atypical sleep period of 8 hours or less. The cardiac support system mayalso include external monitoring devices useful to notify the patientand/or other caregivers of system status and/or other issues.

Implantable Rotary Blood Pump

An implantable rotary blood pump may assist or fully support therequired blood flow of a patient. For example, implantable rotary bloodpumps may have circulatory assist uses including, but not limited to,ventricular assist (right, left and both) and heart replacement. For thepurpose of illustration, a highly efficient blood pump is discussedbelow. However, it should be noted that an implantable rotary blood pumpis in no way limited to the specific implementations discussed below.FIG. 1 is a top view of an illustrative implementation of an implantablerotary blood pump 10 hereinafter referred to as pump 10. Pump 10 isformed from pump housing 15 providing inlet 20 and outlet 25 and motorhousing 35. Pump housing 15 is composed of two or more pieces and may bejoined by welding. However, in other implementations, pump housing 15may be joined by fusing, press fit, threading, screw and elastomericsealing, bonding, fasteners, and/or any other suitable joining method orcombinations of joining methods. Motor housing 35 may be joined to pumphousing 15 by welding, fusing, press fit, threading, screw andelastomeric sealing, bonding, fasteners, and/or any other suitablejoining method or combinations of joining methods. Line A-A passingthrough pump housing 15 indicates the plane from which thecross-sectional view in FIG. 2 is provided.

FIG. 2 is a cross-sectional side view of an illustrative implementationof pump 10. Pump housing 15 provides impeller chamber 30 for impeller75. Impeller chamber 30 has inlet 20 for connection to a fluid sourceand outlet 25 for providing fluid to a desired location. Impellerchamber 30 is sealed and pressure tight to prevent fluid fromentering/exiting impeller chamber 30 from locations other than inlet 20and outlet 25.

Motor housing 35 is attached to pump housing 15 to form a fluid and/orpressure tight chamber for motor 40. While motor housing 35 is shown asa separate component from pump housing 15, in other implementations,pump housing 15 and motor housing 35 may be combined to form a singlecombined housing. A cross-sectional view of an illustrativeimplementation of motor 40 and motor housing 35 of pump 10 is shown inFIG. 7. In particular, motor housing 35 is shown separate from pump 10.Motor 40 is entirely contained between pump housing 15 and motor housing35. A high efficiency electric motor can be utilized, such as anelectric motor with efficiency of about 85% or greater. However, inother implementations, any other suitable driving means can be utilized.Motor 40 provides shaft 45 with hub 50 mounted to shaft 45. Hub 50contains one or more permanent magnets and/or magnetic materials 55.Motor 40 rotates shaft 50 causing permanent magnets 55 placed in hub 50to rotate. In some implementations, a motor with a useful life ofgreater than 10 years is utilized. Further, the motor may utilizehydrodynamic bearings with fluid support provided by a fluid other thanblood.

A cross-sectional view of an illustrative implementation of pump housing15 without impeller 75 is shown in FIG. 6. Pump housing 15 may provide anon-ferromagnetic and/or non-electrically conductive diaphragm 60separating impeller chamber 30 from the chamber housing motor 40.Diaphragm 60 defines cavity 70 providing a region for hub 50 to rotatewithin. Additionally, diaphragm 60 may provide cylindrical bearingsurface 65 for impeller 75 to rotate around with hydrodynamic radialsupport. Impeller 75 includes one or more permanent magnets and/ormagnetic materials 80. Permanent magnets 80 allow impeller 75 to bemagnetically coupled to hub 50. This magnetic coupling allows motor 40to cause impeller 75 to rotate when motor 40 rotates hub 50.

Line B-B passing through pump housing 15 indicates the plane from whichthe cross-section view in FIG. 3 is provided. FIG. 3 is across-sectional top view of an illustrative implementation of pump 10.Impeller 75 is composed of an array of arc shaped segments 90 joined bycentral ring 95. Pump housing 15 has volute 110 feeding the outlet 25.In other implementations, volute 110 could be omitted from pump housing15 and outlet 25 could have any suitable orientation and shape. Pumphousing 15 is designed in a manner where impeller 75, when rotated,pressures and moves fluid received from inlet 20 to outlet 25.

Permanent magnets 55 in hub 50 and permanent magnets 80 in central ring95 of impeller 75 form a magnetic coupling between the impeller 75 andhub 50. In contrast to radial magnetic bearings that are arranged torepel each other, permanent magnets 55 and 80 are arranged so that theyare attracted to each other. In order to minimize radial loads,permanent magnets 55 and 80 provide a minimal magnetic coupling or justenough of a magnetic coupling to rotate impeller 75 under load. Theattractive force of the magnetic coupling of permanent magnets 55 and 80also provides axial restraint of impeller 75. For example, axialmovement of impeller 75 would misalign permanent magnets 55 and 80. Themagnetic forces of permanent magnets 55 and 80 would restrain andre-align the magnets. Because of the magnetic forces caused by permanentmagnets 55 and 80, axial movement of impeller 75 may cause axial forceto be exerted on shaft 45 and hub 50 of motor 40, which is thentransferred to bearing(s) (not shown) of motor 40.

Permanent magnets 80 may be sufficiently small in size that they have noimpact on the main fluid flow paths of impeller 75, thereby allowing thedesign of impeller 75 to focus on fully optimizing pump efficiency.These benefits can allow pumping efficiencies of greater than 50% to beachieved.

Impeller internal surface 100 of central ring 95 is utilized to form ahydrodynamic bearing between cylindrical bearing surface 65 and impellerinternal surface 100. Impeller 75 is configured to rotate withinimpeller chamber 30 with full radial hydrodynamic support from thehydrodynamic bearing formed by cylindrical bearing surface 65 andimpeller internal surface 100. A cross section view of an illustrativeimplementation of impeller 75 is shown in FIG. 5 and an isometric viewof an illustrative implementation of impeller 75 is shown in FIG. 8,which more thoroughly illustrate the hydrodynamic bearing.

Pattern grooves on impeller internal surface 100 of impeller 75 create ahigh pressure zone when impeller 75 is rotated, thereby creating ahydrodynamic bearing. For example, symmetrical herringbone groovescreate a high pressure zone where the two straight lines of the V-shapegrooves meet or the central portion of the symmetrical herringbonegrooves. The pressure created by the pattern grooves on impellerinternal surface 100 acts as a radial stabilizing force for impeller 75when it is rotating concentrically. While the implementation shownprovides symmetrical herringbone grooves on internal surface 100 ofimpeller 75, a variety of different groove patterns may be utilized onimpeller internal surface 100 to provide a hydrodynamic bearing, whichis discussed in detail below. Because low loads are exerted on impeller75, the radial hydrodynamic bearing formed between cylindrical bearingsurface 65 and impeller internal surface 100 can provide stable radialsupport of impeller 75.

Impeller 75 may be an open, pressure balanced type impeller to minimizeaxial thrust. Impeller 75 is considered to be open because there is noendplate on either side of arc shaped segments 90. Further, impeller 75is considered to be pressure balanced because it is designed to minimizeaxial thrust during the rotation of impeller 75. However, other types ofimpellers may be suitable in other implementations. Impeller 75 could beany other suitable blade shape, rotate in the opposite direction, ornon-pressure balanced. For example, other suitable impellers may besemi-open type (i.e. end plate on one side of impeller) or closed type(i.e. end plate on both sides of impeller).

FIG. 4 is a close up cross-sectional view of an area C (see FIG. 2) ofan illustrative implementation of pump 10. The magnetic couplingtransmits torque from shaft 45 of the motor 40 to impeller 75. In theimplementation shown, permanent magnets 55 and 80 are radiallydistributed around hub 50 and impeller 75. The poles of permanentmagnets 55 and 80 are arranged to attract to each other. The attractiveforce of the magnetic coupling of permanent magnets 55 and 80 providesaxial restraint of impeller 75. While permanent magnets 55 and 80 areshown as arc shaped like quadrants of a cylinder, it should berecognized that permanent magnets 55 and 80 may be shaped in a varietyof different manners to provide the magnetic coupling. For example, oneor more ring shaped magnets polarized with arc shaped magnetic regions,square/rectangular shaped, rod shaped, disc shaped, or the like may beutilized. In the magnetic coupling arrangement shown, permanent magnets80 are shown in the internal portion of impeller 75. Internal magneticcouplings, similar to the arrangement shown, can be more efficient thanface or external type magnetic couplings that place the magnets in theblades of an impeller or rotor because they have a smaller diameter andless eddy current losses. Diaphragm 60, intermediate the coupling, isnon-ferromagnetic and/or non-electrically conductive to minimize eddycurrent losses. For example, couplings with non-electrically conductingdiaphragms such as bio-compatible ceramic, glass or the like, wouldexhibit less eddy current losses than those with electrically conductingdiaphragms.

In one implementation, motor 40 is of the brushless DC, sensorless, ironcore type electric motor with fluid dynamic bearings. However, in otherimplementations, any suitable type of motor including one or morefeatures such as, but not limited to, brushed, hall-effect sensored,coreless, and Halbach array or any type of bearing such as ball orbushing may be used. Motor housing 35 may include motor controlcircuitry or be configured to operate with remotely located controlcircuits.

Separating motor 40 from impeller chamber 30 may allow a high efficiencymotor to be utilized. For example, incorporating components into a pumpimpeller to form the rotor of an electric motor may compromise thedesign of the pump impeller resulting in reduced efficiency. Further,designing a rotor and stator that is incorporated into the design of apump may result in an electric motor with large gaps between componentsof the rotor and stator, thereby decreasing the efficiency of the motor.The magnetic coupling arrangements utilized in the implementationsdiscussed herein allow a highly efficient motor design to be utilizedwithout compromising the design of an efficient pump impeller.

FIGS. 9A-9K and 10A-10D illustrate various implementations of patterngrooves that may be implemented on impeller internal surface 100. Asdiscussed previously, impeller internal surface 100 provides ahydrodynamic journal bearing. For example, impeller internal surface 100may utilize patterned grooves. The pattern grooves may be of any typeincluding, but not limited to, half herringbone (FIG. 9A), dual halfherringbone (FIG. 9B), symmetrical herringbone (FIG. 9C), dualsymmetrical herringbone (FIG. 9D), open symmetrical herringbone (FIG.9E), open dual symmetrical herringbone (FIG. 9F), asymmetricalherringbone (FIG. 9G), continuous asymmetrical dual herringbone (FIG.9H), asymmetrical dual herringbone (FIG. 9I), asymmetrical openherringbone (FIG. 9J), asymmetrical open dual herringbone (FIG. 9K), orthe like. Flow inducing pattern grooves, such as half herringbonepatterns and asymmetrical herringbone patterns, have the added benefitof producing a substantial secondary flow, particularly along the axisof impeller rotation between cylindrical bearing surface 65 and impeller75, thereby minimizing stagnant flow between cylindrical bearing surface65 and impeller 75. Because stagnant areas may cause blood clots to formin blood pumps, the secondary flow reduces the chances of blood clotsforming. Further, asymmetrical herringbone patterns have the additionalbenefit over half herringbone patterns in that they provide similarradial stiffness as symmetrical herringbone patterns. As shown in FIG.10A-10D, each of the pattern grooves of internal surface 100 can beshaped in a variety of different manners, such as, but not limited to,rectangular grooves, rectangular grooves with a bevel, semi-circulargrooves, elliptical grooves, or the like. In other implementations,impeller internal surface 100 may also be a plain journal bearingwithout pattern grooves or a multi-lobe shape that creates ahydrodynamic bearing. In alternative implementations, the patterngrooves or multi-lobe shapes may be located on the surface ofcylindrical bearing surface 65 facing impeller 75 rather than impellerinternal surface 100 or the pattern grooves may be located on an outerradial surface of impeller 75 or internal radial surface of pump housing15 facing the impeller 75.

FIG. 11 provides a cross-sectional side view of an illustrativeimplementation of housing 150 for pump 120. Similar to theimplementation shown in FIG. 2, pump 120 provides pump housing 150,impeller 125, shaft 130, hub 132, permanent magnets 135 and 140, motorhousing 142, motor 145, and impeller chamber 160, which all provide asimilar function to the components discussed previously. These commonelements may operate in substantially the same manner as previouslydescribed. The substantial differences in the implementations arediscussed below.

The implementation shown in FIG. 2 provided radial support of impeller75 utilizing a hydrodynamic bearing. However in FIG. 11, in addition toa radial hydrodynamic bearing, one or more external planar surfaces ortop surfaces 165 of impeller 125 include pattern grooves providingpartial axial hydrodynamic support.

FIG. 13 is a close up cross-sectional view of an area D of anillustrative implementation of pump 120. Each arc shaped segment 127 ofimpeller 125 includes one or more pattern grooves on top surfaces 165.The pattern grooves on top surface 165 of impeller 125 and internalsurface 155 of housing 150 form a hydrodynamic bearing providing partialaxial hydrodynamic support that prevents or minimizes contact betweenimpeller 125 and housing 150. The pattern grooves on top surface 165 areconsidered to be interrupted because they are separated by the flowchannels of impeller 125.

Pattern grooves on top surface of impeller 125 may be any suitable typeof grooves including, but not limited to, spiral herringbone and spiralgrooves shown in FIGS. 12A and 12B. FIGS. 14A and 14B respectivelyprovide an isometric view of impeller 125 with spiral herringbone andspiral grooves. The arrangement of the pattern grooves on top surfaces165 is balanced so that instability during rotation of impeller 125 isprevented or minimized. For example, all of the top surfaces 165 havepattern grooves in the implementation shown. However, it should berecognized that in other implementations a balanced arrangement of topsurfaces 165 that have pattern grooves and do not have pattern groovesmay be utilized. A balanced arrangement of top surfaces 165 prevents orminimizes the instability of impeller 125. Examples of balancedarrangements for the implementation shown may include, but are notlimited to, all top surfaces 165 with grooves or three alternating topsurfaces 165 with grooves and three without grooves. Flow inducingpattern grooves, such as spiral and spiral herringbone grooves, have theadded benefit of producing a substantial secondary flow, particularlybetween top surface 165 of impeller 75 and internal surface 155 ofhousing 150. Additionally, various pattern groove types includingsymmetrical, asymmetrical, open, and/or dual groove patterns and variousgroove shapes including rectangular, rectangular with a bevel,semi-circular, and elliptical shown in FIGS. 9A-9K and 10A-10D may beutilized. An additional benefit of the hydrodynamic bearing on topsurface 165 of impeller 125 is that it increases impeller stabilityduring rotation by restraining angular motion along axes normal to theaxis of impeller rotation.

FIG. 15 is a cross-sectional side view of an illustrative implementationof pump 170 with a conically shaped impeller 175. Many of the componentsof pump 170 are substantially similar to the components of thepreviously discussed illustrative implementations. These similarcomponents may operate in substantially the same manner as previouslydescribed. As in the previously discussed implementations, impeller 175is magnetically coupled to shaft 180 of motor 182. Permanent magnets 185and 190 couple motor 182 to impeller 175. However, in the implementationshown, impeller 175 is formed in a generally conical shape. Top surfaces195 of impeller 175 facing internal surface 200 of the pump housing 202are shaped in a manner that provides a hydrodynamic bearing betweenimpeller top surfaces 195 and internal surface 200.

FIG. 17 is a close up cross-sectional view of an area E of anillustrative implementation of pump 170. As in the other implementationspreviously discussed, internal surface 205 of impeller 175 may includepattern grooves for a hydrodynamic bearing providing radial support. Topsurfaces 195 of impeller 175 are angled to provide a generally conicalshaped impeller 175. FIGS. 16A-16E are views of various implementationsof impeller 175. Impeller 175 has multiple blade segments 210 that eachhave a top surface 195. Top surfaces 195 of blade segments 210 may belinear (FIG. 16A), convex (FIG. 16B), or concave (FIG. 16C) surfaces.Additionally, FIGS. 16D-16E are views of impeller 175 with convex andconcave top surfaces 195.

One or more of the top surfaces 195 of impeller 175 may incorporateinterrupted pattern grooves of any type including, but not limited to,spiral or spiral herringbone grooves. For example, the interruptedpattern grooves may be similar to the pattern grooves shown in FIGS. 12Aand 12B. The arrangement of the pattern grooves on top surfaces 195 isbalanced so that instability during rotation of impeller 175 isprevented or minimized. For example, all of the top surfaces 195 havepattern grooves in the implementation shown. However, it should berecognized that in other implementations a balanced arrangement of topsurfaces 195 that have pattern grooves and do not have pattern groovesmay be utilized. Flow inducing pattern grooves, such as spiral andspiral herringbone grooves, have the added benefit of producing asubstantial secondary flow, particularly between top surface 195 ofimpeller 175 and internal surface 200 of pump housing 202. Additionally,various pattern groove types including symmetrical, asymmetrical, open,and/or dual groove patterns and various groove shapes includingrectangular, rectangular with a bevel, semi-circular, and elliptical mayalternatively be utilized as shown in FIGS. 9A-9K and 10A-10D. In someimplementations, top surfaces 195 of impeller 175 do not utilize patterngrooves. For example, the conical shaped impeller 175 may be a pressurebalanced type impeller where the magnetic coupling formed by magnets 185and 190 provides sole axial restraint of impeller 175.

In addition to the axial restraint provided by the magnetic couplingdiscussed previously, the hydrodynamic bearing provided by top surfaces195 of impeller 175 partially restrains axial movement in the directionalong the axis of rotation. Because top surfaces 195 are angled, thehydrodynamic bearing of top surfaces 195 also partially restrains radialmotion of impeller 175. Thus, the hydrodynamic bearing of top surfaces195 provides partial radial and axial support for impeller 175. Thehydrodynamic bearings of top surface 195 and impeller internal surface205 and the partial restraint provided by the magnetic coupling increaseimpeller stability during rotation by restraining axial and radialmotion.

FIG. 18 is a cross-sectional side view of an illustrative implementationof pump housing 215 for pump 212. Many of the components of pump 212 aresubstantially similar to the components of the previously discussedillustrative implementations. These similar components may operate insubstantially the same manner as previously described. As in thepreviously discussed implementations, impeller 220 is magneticallycoupled to shaft 225. Permanent magnets 230 and 235 couple the motor toimpeller 220.

Impeller 220 contains permanent magnets 240 and pump housing 215contains permanent magnets 245, 250 thereby forming a magnetic thrustbearing for minimizing axial movement of impeller 220. Permanent magnets245, 250 in housing 215 may be one or more magnets formed into a ring.FIG. 20 is a close up cross-sectional view of an area H of anillustrative implementation of pump 212. Permanent magnets 240 inimpeller 220 and permanent magnets 245 in the top portion of pumphousing 215 are arranged to provide a repulsive force between impeller220 and pump housing 215. Permanent magnets 240 in impeller 220 andpermanent magnets 250 in the bottom portion of pump housing 215 are alsoarranged to provide a repulsive force between impeller 220 and pumphousing 215. The axial restraint forces generated by magnets 240, 245,250 are significantly greater than the attractive forces generated bythe permanent magnets 230 and 235 and thereby provide sole axial supportwith greater stiffness for impeller 220 during rotation. Magnets 240 inimpeller 220 and magnets 245, 250 in pump housing 215 provide largeaxial restraint forces to allow for increased clearances betweenimpeller 220 and pump housing 215 during rotation. The increasedclearances reduce damage to blood and allow for increased flow throughthe clearances during impeller rotation.

FIG. 19 is a cross sectional top view of an illustrative implementationof pump 212. Magnets 240 are arranged radially around impeller 220. Eachblade segment 255 of impeller 220 may provide an opening/region forreceiving one or more magnets 240. Additionally, in someimplementations, the top and/or bottom surfaces of impeller 220 mayincorporate various pattern groove types including spiral, spiralherringbone, symmetrical, asymmetrical, open, and/or dual groovepatterns. Further, various groove shapes including rectangular,rectangular with a bevel, semi-circular, and elliptical may also beutilized as shown in FIGS. 9A-9K and 10A-10D.

Implantable Power Module and Wireless Power Transfer Subsystem

An implantable power module may provide energy storage to power animplantable rotary blood pump. Some currently available power modulesare worn externally and may require a percutaneous wire penetrating thepatient's skin to power the implanted rotary blood pump. Other availablecardiac support systems that do not require a percutaneous wire mayutilize inductive energy transfer to power the implanted blood pumpwirelessly. These systems may also utilize implanted batteries to powerthe blood pump when inductive energy transfer is not provided. However,due to the high power consumption of the implanted blood pump and/orother components, these systems are only capable of operation for ashort duration using power from implanted batteries that have been fullycharged. For example, the LionHeart LVD-2000 utilizes short rangeinductive charging and is capable of approximately 20 minutes ofoperation using power provided by an implanted battery. In contrast, theimplantable power module described herein is capable of operating theimplanted rotary blood pump for an entire day of awake hours using powerprovided by the energy storage device contained within the power moduleimplanted in the patient. The wireless power transfer subsystemdescribed herein is capable of providing power, without the need forpercutaneous wires, to operate the implanted blood pump andsimultaneously recharge the implanted energy storage device during anormal sleep period of 8 hours or less. Moreover, the wireless powertransfer subsystem is capable of providing power using short rangeinductive energy transfer or mid range energy transfer using magneticresonance coupling (MRC).

FIG. 21 is an isometric view of an illustrative implementation of awireless power subsystem and power module for a cardiac support system.Wireless power subsystem and power module may include transmitting coilassembly 310, receiving coil assembly 315, RF power supply 318, andpower module 319. In some implementations, receiving coil assembly 315and power module 319 may be implanted in a patient. Receiving coilassembly 315 and power module 319 may be provided in the same orseparate hermetically sealed biocompatible housing(s). FIG. 22 is anisometric view of an illustrative implementation of a transmitting coilassembly 310. Transmitting coil assembly 310 may include an excitationcoil 320, transmitting resonant coil 325, mounting plate 327, housing330, and cover 331. FIG. 23 is a front view of an illustrativeimplementation of transmitting resonant coil 325, excitation coil 320,and mounting plate 327. Excitation coil 320 is placed close enough totransmitting resonant coil 325 to be inductively coupled such that whenhigh frequency AC power, such as that from an RF power supply 318 shownin FIG. 21, on the order of 30 KHz-15 MHz is supplied to excitation coil320, this causes transmitting resonant coil 325 to resonate resulting ina local time varying magnetic field. This resonant magnetic fieldinteracts with a resonant coil provided by receiving coil assembly 315as shown in FIG. 24. This resonant magnetic field interaction betweenthe transmitting resonant coil 325 and a resonant coil provided byreceiving coil assembly 315 is referred to as magnetic resonancecoupling.

Magnetic resonance coupling is a phenomenon in which two resonantobjects tuned to the same or similar frequency electromagneticallyexchange energy strongly but interact only weakly with othernon-resonant objects. For example, magnetic resonance coupling may allowenergy to be transferred wirelessly between two resonant coils oversignificant distances, whereas inductive coupling requires the two coilsto be placed close to each other.

FIG. 24 is an isometric view of an illustrative implementation of areceiving coil assembly 315. Receiving coil assembly 315 provides areceiving resonant coil 335, power pick-up coil 340, mounting plate 342,hermetically-sealed biocompatible housing 345, and cover 346. FIG. 25 isa front view of an illustrative implementation of receiving resonantcoil 335, power pick-up coil 340, and mounting plate 342. Excitationcoil 320 and power pick-up coil 340 may be made from a minimal number ofconductor loops, and with any suitable conductor material, such asstranded or solid copper wire, so as not to produce too strong inductivecoupling to their respective resonant coils 325 and 335 and therebyminimize the effect on resonant coil natural frequency and Q factor asdiscussed further below. Housing 345 and cover 346 are made of abiocompatible material. Housing 345 and cover 346 secure and sealreceiving coil assembly 315. Note that in other implementations,receiving resonant coil 335 may be separated from receiving coilassembly 315 as discussed in further detail below.

Transmitting resonant coil 325 and receiving resonant coil 335 aredesigned to have closely matched or identical natural resonantfrequencies as defined by equation 1.

$\begin{matrix}{\omega = \sqrt{\frac{1}{LC}}} & \lbrack 1\rbrack\end{matrix}$

where,

-   -   ω=coil natural resonant frequency (radians)    -   L=coil inductance (Henries)    -   C=coil capacitance (Farads)

By doing so, the magnetic field produced by transmitting resonant coil325 causes receiving resonant coil 335 to strongly resonate also,generating its own local time varying magnetic field, and therebyachieves magnetic resonance coupling between the transmitting andreceiving coils. In such a system, power may be transferred wirelesslyand efficiently through this magnetic resonance coupling over a muchgreater distance than that of currently known traditional inductivecoupling. Power pick-up coil 340 is placed close enough to receivingresonant coil 335 so as to receive energy from receiving resonant coil335 inductively, causing an AC voltage across power pick-up coil 340.This AC voltage can then be rectified to a DC voltage and used to poweran implantable medical device and/or recharge implantable batteries.

The amount of energy that can be transferred to receiving resonant coil335 is proportional to the strength of magnetic field emitted fromtransmitting resonant coil 325. The strength of the magnetic fieldemitted from transmitting resonant coil 325 should be maximized for agiven amount of energy input to excitation coil 320 to optimize systemefficiency and power transfer as well as minimize receiving coilassembly 315 size. This is accomplished by choosing a drive frequency Fthat is closely matched or identical to the natural resonant frequenciesω of transmitting 325 and receiving 335 resonant coils and by increasingresonant coil quality factor Q, given by equation 2:

$\begin{matrix}{Q = {\sqrt{\frac{L}{C}}*\frac{1}{R}}} & \lbrack 2\rbrack\end{matrix}$

where,

-   -   Q=coil quality factor    -   L=coil inductance (Henries)    -   C=coil capacitance (Farads)    -   R=coil AC resistance (Ohms) at resonant frequency ω (radians)

Each resonant coil should have a Q factor sufficiently high in order toprovide reasonably efficient energy transfer. The diameter and placementof excitation coil 320 in relation to transmitting resonant coil 325 canbe a variety of different sizes and arrangements, as the transmittingcoil assembly does not have the same size and space constraints as thereceiving coil assembly. In some implementations, it may be desirable tomake the diameter of excitation coil 320 smaller than transmittingresonant coil 325, such that the natural resonant frequency and Q factorof transmitting resonant coil 325 is minimally affected by excitationcoil 320 when placed within the enclosed volume of transmitting resonantcoil 325, as shown in FIG. 22. However, in other implementations, thediameter of excitation coil 320 may be larger than transmitting resonantcoil 325 and/or excitation coil 320 may be angled or out of plane withtransmitting resonant coil 325 to minimize effects on the naturalresonant frequency and Q factor of transmitting resonant coil 325.

One or more components of the receiving coil assembly may be implantedinto the human body. Thus, it may be desirable to minimize the size ofreceiving resonant coil 335 and/or power pick-up coil 340 to beimplanted. For example, the size of a receiving coil assembly may beminimized by placing power pick-up coil 340 within the enclosed volumeof receiving resonant coil 335. The outer diameter of power pick-up coil340 can be made smaller than the outer diameter of receiving resonantcoil 335, such that the natural resonant frequency and Q factor ofreceiving resonant coil 335 is minimally affected by power pick-up coil340 when placed within the enclosed volume of receiving resonant coil335. This provides an optimum state of system tuning for maximum powertransfer and efficiency while minimizing receiving coil assemblythickness and/or volume. It is important to achieve a receiving coilassembly 315 that is thin and implantable to allow for easy implantationand less noticeable implant site for patient comfort and well being. Forexample, in well tuned systems, receiving coil assembly 315 may be oneinch or less in overall thickness. Note that in some implementations,receiving resonant coil 335 and power pick-up coil 340 may be separatedso that the receiving coil assembly implanted in the patient comprisespower pick-up coil 340 and not receiving resonant coil 335. Such anarrangement would minimize the size of components that are implanted inthe patient, but would require receiving resonant coil 335 to be placednear the location where power pick-up coil 340 is implanted.

As can be seen in equations 1 and 2, the factors affecting the coilquality factor Q are coil inductance, capacitance, AC resistance, andresonant frequency. Specifically, to maximize Q factor, the coilinductance and resonant frequency should be maximized while the coilcapacitance and AC resistance should be minimized. However, as can beseen in equation 1, coil inductance and capacitance must be chosencorrectly to achieve a desired coil natural resonant frequency. For theimplantable wireless power transfer subsystem disclosed herein, thedesired coil natural resonant frequency is between 30 KHz-15 MHz.

One method that can be utilized to increase coil inductance is toprovide more coil turns at larger coil diameters. However, more coilturns and larger coil diameters require longer conductor lengths therebyincreasing coil AC resistance and decreasing the benefit of higherinductance on coil Q factor. Furthermore, conductor lengths greater than1/10^(th) of the resonant frequency wavelength λ may adversely impactperformance due to wave reflections. Additionally, more coil turnsfurther increase coil AC resistance because of proximity effect.Proximity effect is a well known phenomenon in which the local magneticfields of adjacent coil turns cause current flow to be constrained tosmaller and smaller conductor areas as more coil turns are added. Thenet effect is that a decreasing portion of available conductor area isutilized as more coil turns are added. For example, the AC resistance ofa coil with 4 turns can be several times higher than a coil of the sameaverage diameter with only 2 turns, even if the conductor length of the4 turn coil is only twice that of the 2 turn coil.

Another phenomenon that increases coil AC resistance relative to DCresistance is the skin effect. Skin effect is caused by the internalmagnetic fields generated within a single turn of conductor, as opposedto proximity effect caused by multiple conductor turns. Skin effect issimilar to proximity effect in that a decreasing portion of availableconductor area is utilized as AC operating frequency is increased. Thisresults in current flow that is more concentrated at the outer surfacesof a conductor as opposed to the interior portion of a conductor. Thedepth to which most of the current flow is constrained in a conductoroperating at a given AC frequency is known as the skin depth and isgiven by equation 3:

$\begin{matrix}{\delta = \sqrt{\frac{2\rho}{f\; \mu}}} & \lbrack 3\rbrack\end{matrix}$

where, δ=skin depth (meters)

-   -   ρ=resistivity of conductor (Ohm-meters)    -   f=operating frequency (radians)    -   μ=absolute magnetic permeability of conductor (Henries/meter)

Therefore, it can be seen for a conductor of thickness T that is muchthicker than the skin depth δ, most of the conductor is not utilized topass AC current. The ratio of conductor thickness T to skin depth δ isknown as the skin depth ratio. It is clear that increasing conductorthickness T above skin depth δ does little to reduce the AC resistanceof a conductor, but merely increases coil volume and mass. However, italso does not make the skin effect worse.

Notably, it is known in close coupled AC inductive transformer designthat increasing conductor thickness T far above skin depth δ can worsenthe proximity effect substantially, especially as more coil turns areadded. For example, a high skin depth ratio above 2 can cause the ACresistance of an inductive transformer coil to be greater than 10 timeshigher than the same coil with a skin depth ratio of 1 or less,depending on the number of coil turns employed and operating frequency.Therefore, the conductor thickness T used in transmitting 325 andreceiving 335 resonant coils is chosen to produce a skin depth ratio ofless than or equal to 2 to minimize proximity effects, reduce coil ACresistance, and increase coil quality factor Q. Similarly, a skin depthratio less than one may be advantageous. In one implementation, copperor silver foil of a thickness less than 0.020 inches is used. Counterintuitively, thin copper foil produces less AC resistance than thickcopper foil for some of the operating frequencies disclosed herein. Byutilizing a thin conductor, it is believed that a quality factor of 100or greater may be achieved. In experiments using thin copper foil, areceiving resonant coil 335 with a quality factor above 300 for a coilsize 3 inches or less in diameter and 0.5 inches or less in width hasbeen achieved, which would result in a receiving coil assemblysufficiently small to implant. A receiving resonant coil 335 of the sizeabove would then allow the entire receiving coil assembly to be lessthan 1 inch thick. Such a receiving resonant coil 335 may enclose anarea of 7.1 in² or less. Further, the total volume of receiving resonantcoil 335 may be 7.1 in³ or less. Additionally, this may result in atransmitting resonant coil 325 with a quality factor above 600 for acoil size 6 inches or greater in diameter and one inch or less in width.Such a transmitting resonant coil 325 may enclose an area of 28.3 in² ormore. Further, the total volume of transmitting resonant coil 325 may be28.3 in³ or more. Using the foregoing transmitting 325 and receiving 335resonant coil diameters may result in a transmitting/receiving resonantcoil diameter ratio of 2:1 or greater which may allow adequate power tobe transferred over a distance equal to or greater than the diameter ofreceiving resonant coil 335. In experiments, we have achieved adequatepower transfer over distances greater than five times the diameter ofthe receiving coil. Such a system design is uniquely suited forimplantable wireless power systems and methods. Unlike traditionalinductive coupling, such systems and methods may be capable oftransmitting adequate power even when transmitting and receiving coilsare laterally or angularly misaligned to a large extent, such as when apatient is sleeping.

As shown in equation 1, once the inductance of resonant coil 325 or 335is fixed, the proper capacitance must be present for the coil toresonate at a desired frequency ω. Coil capacitance can either beintrinsic, added in the form of a fixed or variable capacitor, or bothintrinsic and added. Intrinsic capacitance is that which is formed bythe coil geometry itself. For example, a coil with turns made fromcopper or silver foil separated by one or more insulating dielectricmaterials such as PTFE, low-loss PTFE, polyethylene, polypropylene,vacuum, an inert gas, or air could be analogous to a flat platecapacitor of equal plate area and plate separation distance. However,intrinsic coil capacitance cannot be calculated in the same manner as aflat plate capacitor due to the effect of multiple turns. Manydielectric materials, such as those listed previously, are suitable toprovide this intrinsic capacitance; however it is important that thematerials have a low dielectric dissipation factor to not detrimentallyimpact the overall coil Q factor. To maintain an overall coil Q factorsufficiently high for adequate power transfer, the one or moreinsulating materials should have a dielectric dissipation factor of 0.01or less at the coil resonant frequency.

It is desirable for transmitting 325 and receiving 335 resonant coils tohave as little intrinsic capacitance as possible, if the intrinsiccapacitance is formed partially or fully by a solid dielectric material.This is done to minimize the temperature sensitivity of the resonantcoils which can shift their resonant frequencies and detune the system,resulting in lost power and efficiency. One method that can be utilizedto assist in stabilizing the resonant frequency of receiving resonantcoil 335 is to maintain receiving resonant coil 335 at a relativelyconstant temperature, such as that provided by implanting inside thehuman body at a temperature of 37+/−5 degrees C. Additionally,transmitting resonant coil 325 may be maintained at a relativelyconstant temperature of 25+/−5 degrees C. with the use of cooling fanscontained in durable housing 330.

FIGS. 26 a and 26 b are front and side views of illustrativeimplementations of a resonant coil, such as transmitting or receivingresonant coil, with single wrap conductive foil. In one implementation,resonant coils 325 and 335 achieve very low intrinsic capacitance usinga flat conductor geometry, such as conductive foil 350 constructed fromone or more high conductivity materials such as copper or silver,separated by an insulating medium 355 composed of one or more lowdielectric constant materials such as PTFE, low-loss PTFE, polyethylene,polypropylene, vacuum, an inert gas, air, or any combination thereofwith relatively large spacing D between turns as shown in FIG. 26 a. Asdescribed previously, the one or more insulating materials should have adielectric dissipation factor of 0.01 or less at the coil resonantfrequency to maintain an overall coil Q factor sufficiently high foradequate power transfer. Spacing D indicates the total thickness of theinsulating medium 355. In some implementations, insulating medium 355may be composed of at least one solid material with a polygonal crosssection that also provides mechanical support for the conductive foil350. A polygonal cross section, defined as a cross sectional shape withall straight sides, is chosen as it is a readily available form of PTFE,low loss PTFE, polyethylene, and polypropylene and results in a volumeefficient resonant coil assembly. In the side view shown in FIG. 26 b,width W may indicate the width of the conductive foil 350 and insulatingmedium 355. The amount of capacitance can be varied byincreasing/decreasing the spacing D between coil turns orincreasing/decreasing the conductor width W. Spacing D can be keptconstant or varied between any adjacent turns so long as it results inthe desired low intrinsic capacitance. One or more fixed or variableexternal capacitors with low temperature sensitivity may be added acrossthe start and end of the coil turns to tune the coil to a desiredresonant frequency. Low dielectric dissipation factor externalcapacitors should be used so that when combined with the insulatingmedium 355, the combined dielectric dissipation factor of the externalcapacitors and insulating medium 355 is low to maintain an overall coilQ factor sufficiently high for adequate power transfer. Low temperaturesensitivity external capacitance with a temperature coefficient of lessthan 3000 ppm/degree C. should be used and the external capacitanceshould be at least one tenth the intrinsic capacitance to positivelyimpact the thermal stability of the overall coil capacitance. The start360 and end 365 of conductive foil 350 may be approximately within 45degrees of each other to minimize external capacitor lead length.

In an illustrative implementation, conductive foil 350 used in resonantcoils 325 and 335 is chosen with a thickness T, such that the skin depthratio is less than 2 for a given operating resonant frequency between 30kHz-15 MHz. This is done to decrease the coil AC resistance and therebyincrease coil Q factor. To further decrease coil resistance, theconductive foil 350 may be provided on both sides of an electricallynon-conductive round or rectangular spiral coil form, made from materialsuch as ABS or polycarbonate. For example, the electricallynon-conductive round or rectangular spiral coil form may be doublewrapped by adhering conductive foil 350 to both the inside and outsidesurfaces of the coil form. This effectively provides two single layersof conductive foil 350 on opposing faces of the non-conductive form,which may have multiple benefits. First, the conductor cross sectionarea is doubled, resulting in lower coil DC resistance and possiblehigher coil Q factor, with only a small increase in coil size and mass.Second, the capacitive spacing D can be formed with an all air, inertgas, or vacuum gap, making the dielectric dissipation factor low and theintrinsic capacitance of the coil very low and inherently temperaturestable. This is beneficial in keeping the system tuned to a desiredresonant frequency for maximum efficiency and power transfer. Conductivefoil 350 may be adhered to the electrically non-conductive form with anysuitable adhesive such as epoxy, urethane, silicone, or acrylic. In someimplementations, conductive foil 350 may also extend over the edges ofthe coil form to make electrical contact between foil on the inside andoutside surfaces of the coil. Alternately, if a coil form with circularcross section is used, conductive foil 350 may be wrapped around theentire circumference of the coil form to eliminate currentconcentrations at conductor edges.

Alternately, the conductive path of resonant coils 325 and 335 may beformed by electroplating or electroless plating of a conductive materialsuch as copper or silver onto a suitable electrically non-conductiveform. This may result in multiple advantages. First, manufacturingmaterial and labor costs may be lower due to eliminating costsassociated with adhering conductive foil to an electricallynon-conductive form. Secondly, the conductive path formed byelectroplating or electroless plating is continuous around theelectrically non-conducting form which may further lower coil ACresistance and increase coil Q factor. The thickness of the conductivelayer plated onto the electrically non-conductive form is chosen suchthat the skin depth ratio is less than 2 for a given operating frequencybetween 30 kHz-15 MHz. Again, this is done to minimize the proximityeffect and lower coil AC resistance and increase coil Q factor.Electroless plating of conductive material onto an electricallynon-conductive form may be preferred over electroplating to produce amore uniform conductor thickness throughout the coil geometry. Theelectrically non-conductive form may be made from a material that isreadily platable with copper or silver such as ABS, nylon, orpolycarbonate.

Another factor which determines how much power can be transferredbetween transmitting coil assembly 310 and receiving coil assembly 315is the coupling coefficient between transmitting 325 and receiving 335resonant coils. The coupling coefficient is a function of coil geometryand varies between 0 and 1. Higher coupling coefficients allow morepower to be transferred between resonant coils across greater distances.Coil turns of transmitting 325 and receiving 335 resonant coils arespaced apart (distance D shown in FIG. 26 a) by at least 0.003 inches,preferably 0.030 inches or greater, to increase the coupling coefficientbetween coils. This also has the added benefit of reducing resonant coilintrinsic capacitance.

An alternate conducting medium for resonant coils 325 and 335 forfrequencies in the range 30 kHz-5 MHz is Litz wire, which is a type ofcable designed to reduce the skin effect and proximity effect losses inconductors, thereby reducing the AC resistance. Litz wire consists ofmultiple conductors in the form of thin round wire strands, individuallyinsulated and twisted or woven together, following one of severalprescribed patterns intended to equalize the proportion of the overalllength over which each strand is at the outside. Preferably, each strandhas a skin depth ratio of approximately one or less for a givenoperating frequency between 30 kHz-5 MHz. Operation in lower frequencyranges, for example, 135 kHz, provides several advantages for use inmedical implants, including, but not limited to, increasedelectromagnetic safety and improved performance in the presence ofmetallic shielding.

Because of the criticality of this wireless power system in life supportapplications, such as a VAD or TAH, fault-tolerance is desired. If afailure were to occur which impairs the power transfer using magneticresonance coupling, the excitation coil 320 and power pick-up coil 340could be used directly as power transfer coils utilizing traditionalinductive coupling over a shorter distance. For example, transmittingcoil assembly 310 may be placed on the patient's body near the locationof receiving coil assembly 315. To minimize the inductive couplingdistance and maximize the power transfer, in some implementations it maybe desirable to orient the excitation coil 320 and the power pick-upcoil 340 proximate to each other with their respective transmitting andreceiving resonant coils 325 and 335 being oriented distally. In otherimplementations, a second excitation coil separate from the transmittingcoil assembly may be used to supply power inductively to the powerpick-up coil 340. A suitable frequency range of operation for thisinductive backup mode is 30 kHz-1 MHz, with an exemplary value being 135kHz. While this backup mode operation is suitable for all of thepreviously described implementations, it is especially well suited forthe Litz wire resonant coil because both the magnetic resonance couplingand the backup inductive coupling may be operated at the same frequency,simplifying system design and reducing complexity. In an alternativefault-tolerance approach, the receiving resonant coil 335 may be removedfrom the receiving coil assembly 315 and used as an external(non-implantable) resonator when placed in proximity to the powerpick-up coil 340 in the receiving coil assembly 315. The power pick-upcoil may then be used for inductive coupling as well as for collectingpower from the external receiving resonant coil when magnetic resonancecoupling is available.

The power transfer efficiency of magnetic resonance coupling isincreased when the Q factor of either or both of the resonant coils 325and 335 is increased. Additional “sympathetic” resonant coils, meaningthose which closely match or are identical to the resonant frequency ofthe transmitting and receiving resonant coils 325 and 335, may be usedto increase the power transfer efficiency and range of the transmittingand receiving resonant coils 325 and 335. For example, one or moresympathetic resonant coils 370 may be placed near the transmittingresonant coil 325 to improve the power transfer efficiency as shown inFIG. 27. The additional coils may be placed in geometric positions thatenhance the directionality or universality of the power transfer. Forexample, the additional coils may be placed at angle(s) relative to thetransmitting resonant coil 325 that increase the spatial coverage of theimplantable wireless power subsystem. In some implementations,additional coils may be placed near or around the receiving resonantcoil 335. The sympathetic resonant coil 370 shown in FIG. 27 isillustrative only; sympathetic resonant coil 370 may be in any shape,form factor, or quantity necessary to enhance the efficiency or range ofpower transfer among the resonant coils 325, 335 and 370. Sympatheticresonant coil 370 should have a Q factor sufficiently high in order toprovide reasonably efficient energy transfer. It may be advantageous toplace one or more sympathetic resonant coils so that they, along withthe transmitting resonant coil 325, are over-coupled. When a firstresonant coil is placed within a critical coupling distance near anotherresonant coil, the resonant coils have a tendency to operate optimallyat a shared resonant frequency different from their independent naturalresonant frequency, which is described as over-coupled. In contrast,when a first resonant coil is substantially distant from anotherresonant coil, or outside of a critical coupling distance, the resonantcoils maintain optimum operation at their respective natural resonantfrequencies, which is described as under-coupled. In such a system, thetransmitting resonant coil 325 and one or more sympathetic resonantcoils 370 produce a magnetic resonance field which shares and storesenergy. This use of sympathetic resonant coil 370 is different from ause which would transfer energy from transmitting resonant coil 325 toreceiving resonant coil 335 via “repeating” or “bucket brigade”architecture wherein sympathetic resonant coil 370 is an intermediary.Instead, this over-coupled mode ensures that the sympathetic resonantcoil 370 has a shared resonant frequency with transmitting resonant coil325. When receiving resonant coil 335 is substantially distant fromtransmitting resonant coil 325 and sympathetic resonant coil 370,receiving resonant coil 335 may be under-coupled. Alternatively, asreceiving resonant coil 335 moves substantially near either transmittingresonant coil 325 or sympathetic resonant coil 370, receiving resonantcoil 335 may be over-coupled and produce a shared resonant frequency.

The hermetically-sealed biocompatible housing 345 and cover 346 arepreferably composed of geometries and materials which do not adverselyaffect the Q factor of the receiving resonant coil 335 or the powertransfer efficiency of the wireless power subsystem. Such materials mayinclude, but are not limited to, polyetheretherketone (PEEK),polyetherimide (ULTEM), polysulfone (UDEL), polytetraflouroethylene(PTFE, Teflon), polyurethane (Tecothane), and silicone. Additionally,the geometries and materials are chosen to provide electrical insulationfor the potential high voltages that may be generated in the receivingresonant coil 335, as well as provide spacing necessary to minimizeadverse impacts on the quality factor Q of receiving resonant coil 335due to extraneous materials. Environmental capacitance, meaningcapacitance in the vicinity of transmitting resonant coil 325 orreceiving resonant coil 335, adversely affects the resonant frequency ofcoil 325 or 335 and consequently must be minimized. Therefore, thehermetically-sealed biocompatible housing 345 and cover 346 providespacing around coil 335 and a stable electrostatic environment intendedto stabilize environmental capacitance. In this way, thehermetically-sealed biocompatible housing 345 and cover 346 minimizedetuning and Q factor reduction which would otherwise occur were housing345 and cover 346 not designed specifically for that advantage. Sealingof biocompatible housing 345 may be accomplished with an enclosedhousing or potting of an open housing any suitable potting compound. Inother implementations, sealing may be accomplished by potting the entireassembly of receiving coil assembly 315. While the hermetically-sealedbiocompatible housing 345 is shown without other electronics ormechanical components common to active implantable medical devices, suchas batteries, power rectification and conditioning circuitry, connectorsand the like, such components may be included in or attached to housing345. In some implementations, such components may be housed in aseparate biocompatible housing. In other implementations, it may beadvantageous to perform AC/DC rectification and some or all DC filteringwithin receiving coil assembly 315 to reduce high frequency losses whichmay occur in the implantable biocompatible cable connecting receivingcoil assembly 315 to the power module 319. In such cases the rectifier375 may be placed adjacent to or inside power pick-up coil 340 as shownfunctionally in FIG. 28. Similarly, some or all DC filter components380, such as capacitors and inductors for a π type filter, may be placedadjacent to or inside power pick-up coil 340 also shown functionally inFIG. 28. One advantage of this approach is that the intrinsiccapacitance and inductance of the implantable biocompatible cableconnecting receiving coil assembly 315 and power module 319 may beleveraged as part of a π filter. All electronic components for thewireless power subsystem can be selected for high reliability. Highreliability is especially desirable for components that are to beimplanted in a patient to avoid surgery to remove or repair the system.Likewise, all components of the subsystem may be selected forcompatibility with the electromagnetic fields which will be producedduring the energy transfer.

The resonant coils 325 and 335 implementation previously described is aright circular spiral coil, where the start and end of the coilconductor is within 45 degrees of each other in order to reduce theeffective antenna dipole and reduce electromagnetic radiation. In otherimplementations, any suitable coil arrangement may be utilized, such asa rectangular coil, a helical coil, a square coil, or any other suitablestructure. The number of turns may be one or more. The coil may becomposed of a solid conductor, hollow conductor, flat conductor, Litzwire, any other suitable conductors, and/or a combination thereof. Allmanner of coil shapes, including, but not limited to, circles, squares,rectangles, octagons, other polygons, regular areas and irregular areas,are within the scope of this invention. While the illustrativeimplementations utilize copper or silver conductor coils, any suitableconductive materials or combination of conductive materials may beutilized.

Cardiac Support System

FIG. 29 is an illustrative implementation of a cardiac support system,which may include an implantable rotary blood pump, an implantable powermodule, a wireless power transfer subsystem, a patient monitor, and aprogrammer. The illustrative implementation shown provides blood pump410, outflow graft 412, pump cable 415, power module 420, receiver cable425, receiving coil assembly 430, transmitting coil assembly 440, radiofrequency (RF) power supply 445, patient monitor 450, and programmer460. Blood pump 410 has a fluid inlet and a fluid outlet connected to animpeller chamber, within which an impeller resides. The impeller isfashioned so that when it is rotated by a force external to theimpeller, it will impart motive energy to a fluid in the impellerchamber to increase flow of that fluid from the inlet to the outlet.When applied to the cardiac circulatory system, the increased flow maybe used for therapeutic purposes such as, but not limited to,ventricular assist (right, left and both) and heart replacement.

Blood pump 410 inlet is preferably attached to the left ventricular apexof the heart directly or via cannula (not shown) and the outlet ispreferably attached to the aorta via outflow graft 412. The impellerchamber is composed of a biocompatible material, such as titanium or anyother suitable material. The impeller chamber further is fashioned tominimize adverse impact to the blood which flows through the impellerchamber.

Blood pump 410 impeller is configured to rotate and impart force onblood moving from the inlet and delivering blood to the outlet. Toinduce blood flow, as with any fluid, power must be imparted to thefluid. The hydraulic power necessary to produce fluid flow is:

$\begin{matrix}{W_{flow} = \frac{{flow} \times {density} \times \Delta \; P \times g}{6 \times 10^{7}}} & \lbrack 4\rbrack\end{matrix}$

where W_(flow) is in watts, flow is the desired fluid flow rate inliters per minute (LPM), density is the fluid density in kg/m³, ΔP isthe differential pressure between inlet and outlet in vertical columnmm, and g is the acceleration of gravity in m/s². The cardiac supportsystem minimizes power losses while providing the power necessary forthe intended blood flow in order to provide an energy-efficient systemsuitable for an unencumbered daily lifestyle. The power necessary torotate an impeller is:

$\begin{matrix}{W_{impeller} = \frac{{torque} \times 2\pi \times {speed}}{60}} & \lbrack 5\rbrack\end{matrix}$

where W_(impeller) is in watts, torque is the turning force of theimpeller in Newton-meters, and speed is the impeller rotational rate inrotations per minute (RPM). The power efficiency of a pump can bedefined as:

$\begin{matrix}{\eta_{pump} = \frac{W_{flow}}{W_{impeller}}} & \lbrack 6\rbrack\end{matrix}$

Blood pump 410 impeller is rotated by a motor, which is magneticallycoupled to the impeller. A significant advantage of the magneticcoupling is that the motor chamber and the impeller chamber remainisolated, thereby avoiding blood damage from the motor and couplingbetween the motor and the impeller. When the chambers are not separated,the motor must be suitable for operating in blood. In other blood pumps,the motor is mechanically coupled to an impeller in a separate chamberthrough seals or the like to separate the blood from the motor. However,the mechanical coupling and seals may result in blood clots, and bloodmay enter the motor if the seals fail. The efficiency of a motor whichconverts electrical power into mechanical rotational power can bedefined as:

$\begin{matrix}{\eta_{motor} = \frac{W_{impeller}}{W_{motor}}} & \lbrack 7\rbrack\end{matrix}$

where W_(motor) is the motor input electrical power in watts, which isproportional to the product of motor input voltage in volts (V_(motor))and motor input current in amps (I_(motor)). The typical maximum motorefficiency is:

$\begin{matrix}{\eta_{motorMax} = \left( {1 - \sqrt{\frac{I_{0}}{I_{A}}}} \right)^{2}} & \lbrack 8\rbrack\end{matrix}$

where I_(o) is the no-load motor current in amps and I_(A) is the motorstall current in amps. The maximum motor efficiency is reached when:

I _(motor)=√{square root over (I ₀ ×I _(A))}

The cardiac support system maximizes motor efficiency by minimizing I₀and maximizing I_(A). To minimize I₀, motor frictional losses areminimized by utilizing low-loss internal bearings. Blood pump 410 mayutilize any suitable motor, such as a brushless DC (BLDC) motor toeliminate frictional losses of brushes and increase reliability. Tomaximize I_(A), motor windings are designed using low-loss conductors,frame materials, geometries to reduce winding losses, and small air gapsto minimize flux losses. During operation, when motor speed adjustmentis needed, this invention may adjust the speed by adjusting V_(motor)(wherein speed is proportional to V_(motor)) while simultaneouslymaintaining maximum motor efficiency by specifically operating the motorat or near I_(motor) equal to the value given in equation 9.

The electrical power and control signals are delivered to the blood pump410 through pump cable 415. Pump cable 415 may connect to blood pump 410and power module 420 using connectors on one or both ends suitable forimplanted medical devices.

By utilizing a power-efficient blood pump 410, the cardiac supportsystem minimizes the amount of power required to operate the system.Note that the cardiac support system utilizes a magnetically coupledmotor and impeller that are provided in separate chambers. Some bloodpumps have attempted to reduce pump size by integrating the impellerinto the motor as the rotor itself. This approach may reduce pump size,but because of design constraints, both the impeller and motor aredifficult to optimize for power efficiency. This approach may result inan impeller with suboptimal efficiency, a motor with suboptimalefficiency, or both. In contrast, a magnetically coupled motor andimpeller allows a highly efficient motor to be utilized withoutaffecting the efficiency of the impeller design. Surprisingly, thismagnetic coupling approach does not result in a pump significantlylarger than other LVADs. More importantly, the highly efficient bloodpump 410 allows energy storage needed to power the pump for a full dayof awake hours to be implanted in a patient. While the LionHeartLVD-2000 uses an implanted rechargeable energy source in the patient,the rechargeable energy source is only capable of providing shortduration operation. Blood pump 410 is capable of operating for an entireday of awake hours on power provided by an implanted rechargeable energysource.

Blood pump 410 is small enough for implantation above the diaphragm orpericardially. Power module 420 is separated from blood pump 410 so thatpower module 420 may be implanted outside of the pericardium. Forexample, power module 420 may be implanted subcutaneously in thepectoral region near the clavicle, such as done with implantablepacemakers and defibrillators. The separated implant location of powermodule 420 is advantageous should the module need to be replaced, inwhich case only subcutaneous outpatient surgery would be needed insteadof surgery requiring pericardial intrusion. Because of the cardiacsupport system energy efficiency, power module 420 volume may beapproximately 150 cc or less, which is small enough for implantationabove the diaphragm or pectorally. Pump cable 415 connects power module420 to blood pump 410, which allows power module 420 to power andcontrol blood pump 410. Pump cable 415 may utilize multi-filar MP-35Nwire, or other biocompatible metals or alloys, fabricated to providehigh reliability and long-term durability. The conductors may beelectrically insulated from each other and from the tissue surroundingthe implanted pump cable 415 utilizing flexible biocompatible materials,such as silicone, Silastic, and/or any suitable material(s).

Power module 420 is contained in a housing composed of a biocompatiblematerial. The housing provides a hermetic seal for the components ofpower module 420. However, it is recognized by one of ordinary skill inthe art that various components of power module 420 may be provided inseparate housings and/or incorporated with other components of thecardiac support system.

Receiving coil assembly 430 may be connected to power module 420 byreceiving cable 425. Receiving coil assembly 430 may be approximately100 cc or less in volume, and is small enough for implantation above thediaphragm or pectorally. Transmitting coil assembly 440 is capable oftransferring electromagnetic energy to receiving coil assembly 430through the patient's body. Transmitting coil assembly 440 and receivingcoil assembly 430 are utilized to power blood pump 410, provide energyto be stored by a rechargeable energy source powering blood pump 410, orboth. Receiving coil assembly 430 and transmitting coil assembly 440 arecapable of magnetic resonance coupling (MRC), which provides asignificantly greater electromagnetic recharging distance than inductivecoupling. Receiving coil assembly 430 may be operated in either MRC orinductive coupling modes, the latter requiring a shorter electromagneticrecharging distance. Monitor 450 may be utilized to monitor blood pump410 and may transmit/receive data via RF or LF electromagnetic coupling.Programmer 460 may be utilized to operate/control blood pump 410 and maytransmit/receive data via RF or LF electromagnetic coupling. In someimplementations, monitor 450 and programmer 460 may be combined into asingle device.

FIG. 30 is an illustrative implementation of power module 420. Powermodule 420 may include an output 517, an input 527, controller 500,motor control 510, system control 520, sensors and conditioning 530,energy receiver manager 540, power manager 550, data communication 560,and rechargeable energy source 570. Rechargeable energy source 570provides the energy necessary to operate blood pump 410 as well asimplantable controller 500. Rechargeable energy source 570 may be a highenergy density rechargeable battery based upon lithium chemistry, suchas Li-ion. However, any suitable energy storage mechanism may be used asrechargeable energy source 570, such as non-chemical rechargeable energystorage devices such as high energy capacitors. Further, battery cell(s)or capacitor(s) may be used in combination. An exemplary energy densityfor rechargeable batteries is at least 150 watt-hours/liter, with Li-ionbatteries exceeding 250 watt-hours/liter. Depending upon the energycapacity needed, rechargeable energy source 570 may be the largestimplantable component in volume. By utilizing high energy densityrechargeable energy source 570, power module 420 is capable of poweringblood pump 410 for a full day of awake hours, while still remaining asize suitable for pectoral implantation.

Power module 420 also contains the electronics provided by controller500. These electronics may be fabricated of one or more integratedcircuits and passive electronic components. Controller 500 may providereprogrammable/re-configurable software, firmware, and/or hardware.

Power module 420 receives power and/or control signals from input 527,which are routed to motor control 510. Motor control 510 operates themotor in blood pump 410 at the peak efficiency described previously.Motor control 510 may use pulse width modulation (PWM) of a DC powersignal to control the speed of blood pump 410. Motor control 410 may usepower transistors (e.g. MOSFETs) to perform the PWM switching, whereinthe power transistors are designed to minimize switching losses. The PWMfrequency may be selected with consideration of inductances andcapacitances in pump cable 415 and blood pump 410 to reduce power lossesdue to reactive mismatches.

One function of power manager 550 is to provide power as needed to thefunctional blocks shown in controller 500. For example, motor control510 may receive power from the rechargeable energy source 570 or energyreceiver manager 540 via power manager 550. Power manager 550 maycondition or modify power characteristics, such as voltage. However, tomaximize efficiency, power manager 550 may route power with little or noconversion to minimize conversion losses as appropriate. Motor control510 may contain sufficient switching logic to efficiently maintain thespeed of blood pump 410 as long as the voltage provided to motor control510 is between 8 to 20 volts. Power manager 550 may then route powerfrom energy receiver manager 540 or rechargeable energy source 570 toblood pump 410, while either of their available voltages is between 8 to20 volts. Alternatively, power manager 550 may convert an availablevoltage to a desired range when it is outside of the desired range. Insome implementations, power manager 550 may utilize fixed highefficiency conversion circuitry to supply ultra-low-wattage power forother circuitry in controller 500.

Motor control 510 may receive control information from system control520 and return operational status, such as drive currents and speedinformation. Motor control 510 may be in power module 420. However, inother implementations, motor control 510 may be placed in blood pump410. The motor, which may be a BLDC motor, may be a multi-phase motorrequiring commutation of the signals delivered to each phase. In suchimplementations, motor control 510 may provide the electricalcommutation, examples of which are trapezoidal or vector-sinusoidal. Thecombined power efficiency of motor control 510 and power manager 550 canbe defined as:

$\begin{matrix}{\eta_{control} = \frac{W_{motor}}{W_{eSource}}} & \lbrack 10\rbrack\end{matrix}$

where W_(eSource) is the power from the rechargeable energy source 570used by motor control 510 and power manager 550 to produce W_(motor).

Controller 500 may also include energy receiver manager 540, which iselectrically and operationally connected to receiving coil assembly 430via receiver cable 425. Energy transferred to receiving coil assembly430 by magnetic resonance coupling or inductive coupling is provided toenergy receiver manager 540 for storage and/or to power blood pump 410.Energy receiver manager 540 may additionally include power conditioningand protection circuitry appropriate for receiving coil assembly 430 tomaximize electromagnetic energy transfer and transfer distance. In someimplementations, receiving coil assembly 430 may be incorporated intopower module 420, eliminating receiver cable 425 and reducing the numberof components that are implanted.

Power manager 550 efficiently provides power to the other components ofpower module 420 and receives that power from rechargeable energy source570 and/or energy receiver manager 540. Power manager 550 is alsoresponsible for recharging rechargeable energy source 570 using powerfrom energy receiver manager 540. While energy receiver manager 540 isintended primarily for recharging rechargeable energy source 570 throughpower manager 550, it may also be used for continuous power delivery toblood pump 410 from motor control 510 through power manager 550.Rechargeable energy source 570 acts primarily as an energy storage andsource for periods when energy receiver manager 540 is not receivingsufficient power to power blood pump 410 through motor control 510.

Power manager 550 is responsible for all power conditioning, includingDC-DC conversion, current limiting, over-power protection, and/ormanagement of rechargeable energy source 570. Management of rechargeableenergy source 570 may include management of one or more battery cellsand charge/discharge operations suitable for long-term preservation andmanagement of the battery cells. Power manager 550 may provide efficientenergy delivery or conversion from rechargeable energy source 570 tovarious components of power module 420 and may deliver energy to motorcontrol 510 to maximize η_(control). Regarding management of batterycells, power manager 550 may also contain logic for constant-current andconstant-voltage delivery for Li-ion cell recharging.

To send and receive data transcutaneously, controller 500 includes apower-efficient data communication 560 for wireless data transferutilizing radio frequency (RF) communication or the like. Datacommunication 560 performs data formatting, error checking, modulation,transmission and reception of data. For example, data communication 560may utilize the industry-standard MICS band for communication. In otherimplementations, data communication 560 may utilize low frequency (LF)inductively-coupled communication. Data sent to or received fromnon-implanted devices may be communicated to or from system control 520or other components of power module 420 as appropriate. Data may be sentby data communication 560 to patient monitor 450 via RF or LFelectromagnetic coupling so that patient monitor 450 may receiveinformation, such as but not limited to, pump or battery status, andpresent that information to the patient or caregiver using one or moreindicators on or associated with patient monitor 450. The indicators mayinclude, but are not limited to, visual, auditory and tactileindicators, with respective examples being lights, beeps and vibrations.Additionally, data may be sent to data communication 560 from programmer460 via RF or LF electromagnetic coupling so that programmer 460 maysend information, such as but not limited to, new pump speed, for use byimplantable controller 500. Through this interface, operationalparameters may be adjusted by a health care professional or, dependingupon the criticality of the parameter, the patient or caregiver.Operational parameter(s) may be used by controller 500 to alteroperational behavior of blood pump 410, controller 500 itself,rechargeable energy source 570, and/or receiving coil 430.

Power module 420 may optionally include sensors and conditioning 530 tosense, amplify, filter and condition as needed one or more parametersfor utilization by controller 500 for monitoring, recording and/oraltering the operation of the controller 500 or blood pump 410. Forexample, an accelerometer and appropriate circuitry within sensors andconditioning 530 may be provided to detect acceleration as an indicationof increased physical activity. The accelerometer sensor may act as aphysiological demand sensor, in response to which controller 500 mayincrease or decrease the speed of blood pump 410 to meet the patient'sphysiologic needs. In some implementations, sensors and conditioning 530may monitor a voltage differential between points which are electricallyconnected to the housings of power module 420 and blood pump 410,providing an intrinsic cardiac activity sensor. Sensors and conditioning530 then may filter and amplify the cardiac activity signal andcontroller 500 may then use the intrinsic cardiac activity to increaseor decrease the speed of blood pump 410 to meet the patient's needs.There may be therapeutic advantages to varying the pump speed cyclicallyto produce pulsed blood flow which mimics normal heart operation.Because of the efficiency of blood pump 410, less motor torque is neededto produce a particular flow rate than other pumps Thus, blood pump 410is suitable for efficiently producing a pulsed blood flow if desired.Furthermore, the pulse rate range may cover the entire physiologicallyobserved range of 20 to 220 pulses per minute. When blood pump 410 isused in conjunction with controller 500 containing sensors andconditioning 530, the pulsed blood flow may be adjusted dynamically inresponse to the previously described sensors for cardiac activity,physiological demand, or both. Such physiologic flow adaptability may beadvantageous because it facilitates an increased ambulatory and activelifestyle. Sensors and conditioning 530 may contain the physiologic orenvironmental sensors themselves; alternatively, the sensors may belocated outside sensors and conditioning 530, outside controller 500, oroutside power module 420 and located as needed.

System control 520 performs supervisory and data exchange functions.System control 520 may include a microprocessor, microcontroller, orreconfigurable hardware core, whether implemented discretely or within amulti-function integrated circuit. In addition to the power utilized byblood pump 410, components of power module 420 consume a small, butfinite, amount of power independent of the blood flow rate. The averageamount of such power consumed by power module 420 is defined asW_(baseline), and the components of power module 420 are designed tominimize this value. For example, power module 420 may includelow-voltage, low-switching-speed CMOS circuitry and idle power-downtechniques between time-critical operations. Wherever applicable,circuits which consume power during state transitions are operated tominimize state transitions, and circuits which consume power duringstasis are powered down whenever possible.

The cardiac support system implements an energy-efficient system havinga rechargeable energy source 570 capacity requirement of:

$\begin{matrix}{{Energy} = {\left( {\frac{W_{flow}}{\eta_{pump} \times \eta_{motor} \times \eta_{control}} + W_{baseline}} \right) \times T}} & \lbrack 11\rbrack\end{matrix}$

where Energy is in Watt-hours and T is the period of time in hours thatrechargeable energy source 570 is to provide the desired flow. Thecardiac support system, by maximizing the component 11 efficiencies(each of which will be 1 or less) and minimizing W_(baseline), resultsin a reduced-size implantable system capable of providing a desired flowfor a desired period of time between recharges. By maximizing andcombining the component 11 efficiencies, the cardiac support system hasachieved a η_(pump)×η_(motor)×η_(control) value of approximately 0.6.Similarly, this invention has minimized W_(baseline) to a comparativelynegligible level. In contrast, other systems such as those describedpreviously have equivalent η_(pump)>η_(motor)η_(control) values lessthan 0.1.

In order to quantify performance relative to the amount of powerconsumed, an analogous and proportionate summary value for energyefficiency may be defined wherein the rechargeable energy source and allbut one energy output variables are set to clinically relevant standardvalues. An Energy Conversion Ratio (ECR) may be defined as:

ECR=the sustained flow rate (in LPM) a cardiac support system canprovide against 100 mm-Hg differential pressure for 24 hours from a 40Watt-hour rechargeable energy source

To further standardize ECR, a blood hematocrit value of 42% is used,which is the average between male and female patients. Other cardiacsupport systems such as those described previously have ECR values below0.8. The cardiac support system as described herein is capable ofproducing blood flow with an ECR of approximately 4.0.

The invention is not limited to the preferred implementations, butinstead these and other variations of construction are obvious to thoseskilled in the art and are understood to be within the scope and spiritof the invention herein described.

Implementations described herein are included to demonstrate particularaspects of the present disclosure. It should be appreciated by those ofskill in the art that the implementations described herein merelyrepresent exemplary implementation of the disclosure. Those of ordinaryskill in the art should, in light of the present disclosure, appreciatethat many changes can be made in the specific implementations describedand still obtain a like or similar result without departing from thespirit and scope of the present disclosure. From the foregoingdescription, one of ordinary skill in the art can easily ascertain theessential characteristics of this disclosure, and without departing fromthe spirit and scope thereof, can make various changes and modificationsto adapt the disclosure to various usages and conditions. Theimplementations described hereinabove are meant to be illustrative onlyand should not be taken as limiting of the scope of the disclosure.

1-33. (canceled)
 34. A cardiac support system comprising: a rotary bloodpump implantable in a human body and configured to generate an amount ofblood flow; an implantable housing containing (i) a power module coupledto the rotary blood pump and configured to store electrical energyutilized to operate the rotary blood pump and (ii) a receiving coilassembly coupled to the power module; and a transmitting coil assemblyconfigured to be magnetic resonance coupled to the receiving coilassembly and configured to electromagnetically transfer the electricalenergy from the transmitting coil assembly to the receiving coilassembly.
 35. The system of claim 34, wherein the receiving coilassembly comprises a receiving resonant coil having a quality factorabove
 300. 36. The system of claim 34, wherein the transmitting coilassembly comprises a transmitting resonant coil having a quality factorabove
 600. 37. The system of claim 34, wherein the amount of blood flowis generated with an ECR of 1.0 or greater using the electrical energystored by the power module.
 38. The system of claim 34, wherein theamount of blood flow is generated with an ECR of 2.5 or greater usingthe electrical energy stored by the power module.
 39. The system ofclaim 34, wherein the amount of blood flow is generated with an ECR of4.0 or greater using the electrical energy stored by the power module.40. The system of claim 34, wherein the receiving coil assembly is 100cc or less in volume.
 41. The system of claim 34, wherein the powermodule is 150 cc or less in volume.
 42. The system of claim 34, whereinthe power module is configured to store the electrical energy in arechargeable battery or a capacitor.
 43. The system of claim 34, whereinthe implantable housing further contains a controller configured tocontrol a speed of the rotary blood pump.
 44. The system of claim 43,further comprising a cardiac activity sensor or a physiological demandsensor, wherein the controller is configured to adjust the speed of therotary blood pump in accordance with a detected cardiac activity or adetected physiological demand.
 45. The system of claim 43, furthercomprising: a patient monitor wirelessly coupled to the controller,wherein the patient monitor is configured to provide status informationon the rotary blood pump; and a programmer wirelessly coupled to thecontroller, wherein the programmer is configured to change operatingparameters of the rotary blood pump.
 46. The system of claim 34, whereinthe rotary blood pump is configured to produce a pulsatile blood flow.47. The system of claim 34, wherein the implantable housing ishermetically sealed.
 48. A method for treating heart failure, the methodcomprising: generating an amount of blood flow with a rotary blood pumpimplanted in a patient; coupling a transmitting coil assembly to areceiving coil assembly, implanted in the patient, using magneticresonance coupling, such that the transmitting coil assemblyelectromagnetically transfers energy to the receiving coil assembly; andstoring, in a power module implanted in the patient, the energytransferred to the receiving coil assembly.
 49. The method of claim 48,wherein the receiving coil assembly comprises a receiving resonant coilhaving a quality factor above
 300. 50. The method of claim 48, whereinthe transmitting coil assembly comprises a transmitting resonant coilhaving a quality factor above
 600. 51. The method of claim 48, whereinthe amount of blood flow is generated with an ECR of 1.0 or greaterusing the energy stored in the power module.
 52. The method of claim 48,wherein the amount of blood flow is generated with an ECR of 2.5 orgreater using the energy stored in the power module.
 53. The method ofclaim 48, wherein the amount of blood flow is generated with an ECR of4.0 or greater using the energy stored in the power module.
 54. Themethod of claim 48, wherein the receiving coil assembly is 100 cc orless in volume.
 55. The method of claim 48, wherein the power module is150 cc or less in volume.
 56. The method of claim 48, further comprisingdetecting cardiac activity, wherein the amount of blood flow is adjustedin accordance with the cardiac activity detected.
 57. The method ofclaim 48, further comprising detecting physiological demand, wherein theamount of blood flow is adjusted in accordance with the physiologicaldemand detected.
 58. The method of claim 48, further comprisingmonitoring the rotary blood pump wirelessly, wherein status informationis provided visually, acoustically, or tactilely.
 59. The method ofclaim 48, wherein the amount of blood flow generated is pulsatile. 60.The method of claim 48, further comprising programming the rotary bloodpump wirelessly, wherein programming data is provided via RF or lowfrequency electromagnetic coupling.
 61. The method of claim 60, whereinthe programming data alters an operational setting of the rotary bloodpump, power module, or receiving coil assembly.
 62. A cardiac supportsystem comprising: a rotary blood pump implantable in a human body andconfigured to generate an amount of blood flow; and an implantablehousing containing (i) a power module coupled to the rotary blood pumpand (ii) a receiving coil assembly configured to receive energywirelessly and transfer the energy to the power module, wherein theamount of blood flow is generated with an energy conversion ratio (ECR)of 1.0 or greater using the energy in the power module.
 63. The systemof claim 62, wherein the amount of blood flow is generated with an ECRof 2.5 or greater using the energy in the power module.
 64. The systemof claim 62, wherein the amount of blood flow is generated with an ECRof 4.0 or greater using the energy in the power module.
 65. The systemof claim 62, wherein the receiving coil assembly is configured to beinductively coupled to a transmitting coil assembly, wherein thetransmitting coil assembly is configured to transfer the energywirelessly to the receiving coil assembly by induction.
 66. The systemof claim 62, wherein the receiving coil assembly comprises a receivingresonant coil that is configured to be magnetic resonance coupled to atransmitting coil assembly, wherein the transmitting coil assembly isconfigured to transfer the energy wirelessly to the receiving coilassembly by magnetic resonance coupling.
 67. The system of claim 66,wherein the receiving coil assembly comprises a receiving resonant coilhaving a quality factor above
 300. 68. The system of claim 66, whereinthe transmitting coil assembly comprises a transmitting resonant coilhaving a quality factor above
 600. 69. The system of claim 62, whereinthe implantable housing is hermetically sealed.
 70. A method fortreating heart failure, the method comprising: receiving energywirelessly in a receiving coil assembly implanted in a human body;storing the energy in a power module implanted in the human body; andproducing an amount of blood flow from a rotary blood pump, implanted inthe human body, such that the amount of blood flow is generated with anEnergy Conversion Ratio (ECR) of 1.0 or greater using the energy storedin the power module.
 71. The method of claim 70, wherein the amount ofblood flow is generated with an ECR of 2.5 or greater using the energystored in the power module.
 72. The method of claim 70, wherein theamount of blood flow is generated with an ECR of 4.0 or greater usingthe energy stored in the power module.
 73. The method of claim 70,wherein the receiving coil assembly receives the energy wirelessly viainduction.
 74. The method of claim 70, wherein the receiving coilassembly receives the energy wirelessly via magnetic resonance coupling.75. The method of claim 74, wherein the receiving coil assemblycomprises a receiving resonant coil having a quality factor above 300.76. The method of claim 74, wherein the receiving coil assembly receivesthe energy from a transmitting coil assembly comprising a transmittingresonant coil having a quality factor above
 600. 77. A cardiac supportsystem comprising: a rotary blood pump implantable in a human body andconfigured to generate an amount of blood flow; a power module coupledto the rotary blood pump and configured to store electrical energyutilized to operate the rotary blood pump, wherein the power module isimplantable into the human body; a receiving coil assembly coupled tothe power module, being implantable into the human body, and comprisinga receiving resonant coil having a quality factor above 300; and atransmitting coil assembly configured to be magnetic resonance coupledto the receiving coil assembly and configured to electromagneticallytransfer the electrical energy from the transmitting coil assembly tothe receiving coil assembly.
 78. The system of claim 77, wherein theamount of blood flow is generated with an ECR of 1.0 or greater usingthe electrical energy stored by the power module.
 79. The system ofclaim 77, wherein the amount of blood flow is generated with an ECR of2.5 or greater using the electrical energy stored by the power module.80. The system of claim 77, wherein the amount of blood flow isgenerated with an ECR of 4.0 or greater using the electrical energystored by the power module.
 81. The system of claim 77, wherein thereceiving coil assembly is 100 cc or less in volume.
 82. The system ofclaim 77, wherein the power module is 150 cc or less in volume.
 83. Thesystem of claim 77, wherein the power module is configured to store theelectrical energy in a rechargeable battery or a capacitor.
 84. Thesystem of claim 77, further comprising a controller that is implantableinto said human body and configured to control a speed of the rotaryblood pump.
 85. The system of claim 84, further comprising a cardiacactivity sensor or a physiological demand sensor, wherein the controlleris configured to adjust the speed of the rotary blood pump in accordancewith a detected cardiac activity or a detected physiological demand. 86.The system of claim 84, further comprising: a patient monitor wirelesslycoupled to the controller, wherein the patient monitor is configured toprovide status information on the rotary blood pump; and a programmerwirelessly coupled to the controller, wherein the programmer isconfigured to change operating parameters of the rotary blood pump. 87.The system of claim 77, wherein the rotary blood pump is configured toproduce a pulsatile blood flow.
 88. A cardiac support system comprising:a rotary blood pump implantable in a human body and configured togenerate an amount of blood flow; a power module coupled to the rotaryblood pump and configured to store electrical energy utilized to operatethe rotary blood pump, wherein the power module is implantable into thehuman body; a receiving coil assembly coupled to the power module andbeing implantable into the human body; and a transmitting coil assemblyconfigured to be magnetic resonance coupled to the receiving coilassembly and configured to electromagnetically transfer the electricalenergy from the transmitting coil assembly to the receiving coilassembly, wherein the transmitting coil assembly comprises atransmitting resonant coil having a quality factor above
 600. 89. Thesystem of claim 88, wherein the amount of blood flow is generated withan ECR of 1.0 or greater using the electrical energy stored by the powermodule.
 90. The system of claim 88, wherein the amount of blood flow isgenerated with an ECR of 2.5 or greater using the electrical energystored by the power module.
 91. The system of claim 88, wherein theamount of blood flow is generated with an ECR of 4.0 or greater usingthe electrical energy stored by the power module.
 92. The system ofclaim 88, wherein the receiving coil assembly is 100 cc or less involume.
 93. The system of claim 88, wherein the power module is 150 ccor less in volume.
 94. The system of claim 88, wherein the power moduleis configured to store the electrical energy in a rechargeable batteryor a capacitor.
 95. The system of claim 88, further comprising acontroller that is implantable into said human body and configured tocontrol a speed of the rotary blood pump.
 96. The system of claim 95,further comprising a cardiac activity sensor or a physiological demandsensor, wherein the controller is configured to adjust the speed of therotary blood pump in accordance with a detected cardiac activity or adetected physiological demand.
 97. The system of claim 95, furthercomprising: a patient monitor wirelessly coupled to the controller,wherein the patient monitor is configured to provide status informationon the rotary blood pump; and a programmer wirelessly coupled to thecontroller, wherein the programmer is configured to change operatingparameters of the rotary blood pump.
 98. The system of claim 88, whereinthe rotary blood pump is configured to produce a pulsatile blood flow.